Not applicable.
Not applicable.
The present invention relates to PET scanners generally and specifically to a method and apparatus for adjusting PMT gains to compensate for drift due to various operating phenomenon.
Positrons are positively charged electrons which are emitted by radionuclides which have been prepared using a cyclotron or other device. The radionuclides most often employed in diagnostic imaging are fluorine-18, carbon-11, nitrogen-13 and oxygen-15. Radionuclides are employed as radioactive tracers called xe2x80x9cradiopharmaceuticalsxe2x80x9d by incorporating them into substances such as glucose or carbon dioxide. One common use for radiopharmaceuticals is in the medical imaging field.
To use a radiopharmaceutical in imaging, the radiopharmaceutical is injected into a patient and accumulates in an organ, vessel or the like, which is to be imaged. It is known that specific radiopharmaceuticals become concentrated within certain organs or, in the case of a vessel, that specific radiopharmeceuticals will not be absorbed by a vessel wall. The process of concentrating often involves processes such as glucose metabolism, fatty acid metabolism and protein synthesis. Hereinafter, in the interest of simplifying this explanation, an organ to be imaged will be referred to generally as an xe2x80x9corgan of interestxe2x80x9d and prior art and the invention will be described with respect to a hypothetical organ of interest.
After a radiopharmaceutical becomes concentrated within an organ of interest and while the radionuclides decay, the radionuclides emit positrons. Each positron travels a very short distance before it encounters an electron and, when the positron encounters an electron, the positron is annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to medical imaging and particularly to medical imaging using photon emission tomography (PET). First, each gamma ray has an energy of essentially 511 keV upon annihilation. Second, the two gamma rays are directed in substantially opposite directions.
In PET imaging, if the general locations of annihilations can be identified in three dimensions, the shape of an organ of interest can be reconstructed for observation. To detect annihilation locations, a PET scanner is employed. An exemplary PET scanner includes one or more rings of detector modules and a processor which, among other things, includes coincidence detection circuitry. The detector modules are arranged about an imaging area. An exemplary detector module includes six adjacent detector blocks. An exemplary detector block includes an array of 36 bismuth germinate (BGO) scintillation crystals arranged in a 6xc3x976 matrix and four photo-multiplier tubes (PMTs) arranged in a 2xc3x972 matrix to the side of the crystal matrix opposite an imaging area.
When a photon impacts a crystal, the crystal generates light which is detected by the PMTs. The PMT signal intensities are combined to generate a combined analog signal which is converted into a digital signal. For the purposes of this explanation, it will be assumed that the digital value, also referred to as a target value, corresponding to 511 keV is 180. The combined digital signal is compared to a range of values about 511 keV. When the combined signal is within the range, an event detection pulse (EDP) is generated which is provided to the processor coincidence circuitry. In addition, acquisition circuits determine which crystal within a block absorbed the photon by comparing the relative strengths of the PMT signals.
The coincidence circuitry identifies essentially simultaneous EDP pairs which correspond to crystals which are generally on opposite sides of the imaging area. Thus, a simultaneous pulse pair indicates that an annihilation has occurred somewhere on a straight line between an associated pair of crystals. Over an acquisition period of a few minutes, millions of annihilations are recorded, each annihilation associated with a unique crystal pair. After an acquisition period, recorded annihilation data is used via any of several different well known procedures to construct a three dimensional image of the organ of interest.
While operation of a PET detector is relatively simple in theory, unfortunately, despite efforts to manufacture components that operate in an ideal fashion, there is an appreciable variation in how similar detector components respond to identical stimuli. For example, given a detector block including 36 crystals and four PMTs and given the same stimuli, crystals that are positioned proximate the center of the PMT array will typically generate a higher energy value than edge or corner crystals (i.e., crystals that are positioned along the edge of the array or at the corner of the array). This disparate and position dependent energy spectrum occurs because, typically, some of the light generated by an edge or corner crystal is not detected by the PMTs in a single block.
As one other example, even within a single crystal, impacting photons may not generate the same PMT output for various reasons. For instance, some photons are completely absorbed by a crystal while others are not. Completely absorbed photons generate light corresponding to 511 keV while partially absorbed photons generate less than the 511 keV. As another instance, first and second photons may be partially absorbed essentially simultaneously by first and second crystals in the same block. While each photon would be identified if they had been absorbed consecutively, upon simultaneous absorption, the combined energy may erroneously be attributed to a single absorbed photon. In this case detection circuitry may erroneously identify a third crystal between the first and second crystals as the detecting or absorbing crystal.
Thus, while each detected photon should ideally generate a signal having an energy level of 511 keV, in many cases detected photons generate much less energy. For this reason, the energy range used to determine if a combined digital PMT signal corresponds to a detected photon typically is assigned a relatively low threshold value. For instance, in an exemplary PET system the low end of the energy range may be a digital value of 35 corresponding to approximately 100 keV (i.e., any absorption even having an energy greater than 100 keV is assumed to correspond to a photon).
In addition to the potential errors described above, other sources of system error also occur. For instance, given two PMTs and identical stimuli (i.e., input light), a first PMT will typically generate a slightly different output signal than the second PMT. Exacerbating matters, over time PMT performance has been known to degrade due to aging related changes in structure. Further exacerbating matters, PMTs often operate differently when exposed to different operating parameters. For instance, PMT output signals have been known to vary as a function of temperature, ambient magnetic fields and other parameters that are relatively difficult or expensive to control.
To compensate for PMT construction and operating variances, the PET industry has developed various commissioning/calibration procedures and associated hardware and software. Generally, during a calibration procedure, a PET source having a known intensity is provided inside the PET imaging area and PMT signals generated thereby are collected. The collected PMT signals are compared to expected PMT signals and, where there is a difference between the collected and known signals, PMT gains are adjusted to compensate for the differences.
While calibration techniques like the one described above are useful, unfortunately, most calibration techniques require acquisition of massive amounts of data and hence an appreciable amount of time to complete. In addition, many calibration techniques include at least some manual steps that have to be performed by skilled technicians.
Because of the time and skills required to calibrate a PET system, in many cases, calibration will only be performed when absolutely necessary such as after image artifacts begin to appear in generated images. In other cases calibration is performed routinely whether or not the calibration is necessary. For instance, in some cases calibration is performed on a weekly basis. In the case of mobile PET systems (i.e., truck based systems), the system environment and, in particular, ambient magnetic fields, may change on a daily basis. In these cases calibration will typically be performed on a daily basis.
Thus, in some cases where calibration should be performed, calibration may be foregone until a later time while in other cases, where calibration is not necessary, a routine calibration procedure may be performed. In addition, in cases where calibration is only performed when a radiologist begins to recognize artifacts, the radiologist is routinely faced with the question of whether or not to recalibrate.
An exemplary embodiment of the invention includes a method for calibrating PET detector PMT gains in a detector unit including at least one detector block, where a block includes a two dimensional crystal array including crystals arranged adjacent an imaging area and a PMT array including a two dimensional array of PMTs arranged adjacent the crystal array opposite the imaging area, a target energy level being associated with the known average energy of a photon, the method comprising the steps of providing a calibration photon source adjacent the at least one block during a calibration period. For each unit crystal the method further includes obtaining a calibration energy spectrum where the calibration spectrum indicates the number of detected photons at each of several possible energy levels and mathematically combining the calibration spectrum and a crystal specific gain factor to generate a shifted spectrum for the crystal. Thereafter, the method includes combining the shifted spectrums for all unit crystals to generate a unit spectrum, identifying a peak unit energy level for the unit spectrum where the peak unit energy level is the energy level at which the greatest number of photons was detected, comparing the peak unit energy level and the target energy level, based on the difference between the peak unit energy level and the target energy level, adjusting the PMT gains for the unit PMTs.
In at least one embodiment the method further includes the steps of, prior to the step of providing and during a commissioning procedure, providing a commissioning photon source adjacent the at least one block and during a commissioning period, for each unit crystal, obtaining a commissioning energy spectrum where the commissioning spectrum indicates the number of detected photons at each of several possible energy levels, identifying a peak energy level for the commissioning spectrum and mathematically combining the target energy level and the peak energy level to generate the crystal specific gain factor.
The step of mathematically combining to generate the gain factor may includes the step of dividing the target energy level by the peak energy level. Similarly, the step of mathematically combining to generate the shifted energy spectrum may include the step of multiplying each energy level within the spectrum by the crystal gain factor thereby shifting each of the energy level counts.
In addition to the method, the invention includes other similar methods and also contemplates an apparatus that includes either dedicated hardware or that may be implemented in software as computer programs that represent algorithms for execution by a conventional-type digital processor adapted for imaging applications.
These and other aspects of the invention will become apparent from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention and reference is made therefore, to the claims herein for interpreting the scope of the invention.